Ultrasound (US) imaging has provided useful information about the interior characteristics (e.g., organ tissue, material flow, etc.) of a subject under examination. A general US system includes a probe (with a transducer array) that interfaces with a console, which controls the transducer elements of the transducer array to transmit an ultrasonic beam and receive echoes produced in response thereto, which are processed to generate an image(s) of the interior characteristics. The detail and contrast resolutions of the imaging system depend at least on the shape of the ultrasonic beam, which has dimensions both in the imaging plane (azimuth or lateral) and across the imaging plane (elevation).
A one dimensional (1D) transducer array includes a single row of transducer elements arranged along the lateral direction, and the beam is electronically controlled in the lateral direction. The width of the transducer elements is on the order of a wave length. By controlling the delays and weight coefficients in the beamforming, the focus can be controllably moved along a line. In the elevation direction, the height has been several millimeters (e.g., 4 to 20 mm). The focusing in the elevation plane is achieved with acoustic lenses, and the focus is generally fixed. The beam is narrowest at the elevation focus and diverges beyond it. Close to the transducer, the beam is as wide as the transducer array, and away from the elevation focus, the beam becomes even wider.
A 1.5D array has several rows of elements. The effective size of the elements in elevation direction is usually much larger than the width. The outer rows are electrically connected to the middle row. A switch alternately couples outer rows to the middle row, depending on the distance from the transducer surface, creating large elements at large depths. Such arrays have had acoustic lenses that focus the beam in elevation direction. Unfortunately, there is no control over the delays in the elevation plane so there is a trade-off between beam size and the uniformity in the elevation plane. 1.75D array is similar to a 1.5D array, but each element is connected to a channel. This allows electronic focusing in the elevation direction. Unfortunately, the number of channels increases, e.g., from N to 2N, relative to a 1.5D array with N channels.
A synthetic transmit aperture has been used to increase image quality. In one instance, this includes sequentially actuating two or more of the transduce elements, invoking transmissions of two or more ultrasound signals, where the echoes generated in response to each transmission have different phase and/or amplitude information. For each transmission, all of the transducer elements receive echoes, which are beamformed to generate a lower resolution image for each set of received echoes. The lower resolution images are accumulated and/or otherwise combined to generate a higher resolution image. Generally, a higher number of transmissions results in higher image quality, but lower frame rate. Therefore, unfortunately, there is a trade-off between image quality and frame rate.
Coded excitation has been used to increase the signal-to-noise ratio. Examples of spread codes include, but are not limited to, Barker codes, Golay codes, and frequency modulated (FM) pulses. FM modulated pulses tend to be robust to frequency-dependent attenuation and, in many cases, gives the greatest increase in signal-to-noise ratio. An artifact of using FM pulses is the existence of range side-lobes (along the imaging direction). These range side lobes are attenuated by tapering the rising and falling edges of the FM pulse. Typically, a Tukey windowing function is used. This means that the transmitted pulses are both frequency and amplitude modulated. Sending such pulses usually requires either a multi-level linear sender (e.g. 12-bit) or bipolar square wave ([−1, 0, 1]) operating at over 200 MHz clock frequency. Unfortunately, such transmitters tend to be costly.
Obese patients, generally, have a thicker layer of subcutaneous adipose tissue, relative to non-obese patients. The speed of sound in adipose tissue is on the order of 1450 m/s, while the speed of sound in organ tissue tends to be higher. For example, the speed of sound in liver tissue is on average about 1540 m/s. The sound waves refract during their propagation (Snell's law). Delay calculations for beamforming have been based on straight lines of propagation. Unfortunately, this is not an accurate assumption in the case of layered media including adipose tissue and organ tissue.